Throughout this application various publications are referred to in parenthesis. Full citations for these references may be found at the end of the specification immediately preceding the claims. The disclosures of these publications are hereby incorporated by reference in their entireties into the subject application to more fully describe the art to which the subject application pertains.
Magnetic resonance imaging (MRI) is widely used in clinical diagnosis. The SNR of a magnetic resonance image is dependent on a number of factors. A common way to improve SNR from restricted regions is by using a transmit-only volume coil with a receive-only surface coil phased array (1). Since their introduction by Roemer and colleagues (1), phased arrays have proven to be effective in extending the high sensitivity of a single surface coil to a much larger field-of-view (FOV). Due to strong coupling between surface coils in the phased array and the volume transmit coil, the volume coil is used as a “transmit-only” coil, typically being detuned during reception. Although phased arrays have superior SNR near the surface coils, it has not been possible to achieve substantial SNR increases in comparison to optimized head-size quadrature volume coils from central regions (7-8 cm depth) of the human head (2, 3). The region of increased SNR afforded by the arrays can be extended to include the center of the brain if the signals from an optimized volume coil can be combined with those from the array during simultaneous reception. However this combination can only yield significant improvements if 1) the interaction between the volume coil and the surface coil does not substantially degrade the performance of either coil and 2) if the noise acquired by each coil is uncorrelated.
Previously, Kocharian and colleagues employed simultaneous reception with conventional single-loop surface coils using a body coil (4) and a head volume coil (5). The authors did not evaluate quantitatively the change in the SNR due to simultaneous reception as compared with the surface coils alone. However, mutual inductive coupling between the volume coil and the single-loop surface coil altered the RF magnetic field B1 profiles of both coils (5), which would have substantially decreased the overall array performance (6). Furthermore, as the size of the volume coil was reduced to be optimal for heads (5), the inductive coupling between the volume coil and the surface coils increased, thereby further reducing the SNR benefits of the array. Thus, in general single-turn surface coils when used in combination with volumes coils for simultaneous reception yield no or minimal increases in SNR.
To avoid interactions between the volume and the surface coils, Hyde and colleagues have reported a counter rotating current (CRC) surface coil consisting of two parallel rings carrying opposite currents (7). The opposing currents provide intrinsic isolation between the surface coil and the volume transmit coil, which enables simultaneous reception by both coils with improved SNR in areas where they have similar sensitivities. Although the CRC coil has much lower sensitivity than a single-turn coil of the same size when unloaded, when loaded such that sample losses dominate, the CRC's sensitivity becomes virtually indistinguishable from a single-turn surface coil at distances greater than approximately the distance between the two loops of the CRC coil (7). Simultaneous reception using a single CRC coil and a transmit volume coil has been demonstrated previously (8, 9); however, prior to the present invention, the configuration has not been extended to multiple CRC coils in phased arrays. Gradiometer coils, also consisting of two loops with opposite currents, have been described and used previously for nuclear quadrupole resonance (NQR) and low-field magnetic MRI applications (10, 11) to reduce the injected noise.
At higher fields (4 T and above), the sensitivity profile of a head-sized volume coil can be substantially altered due to RF field/tissue interaction (26-28). A very characteristic pattern with the RF magnetic field enhanced at the center of a human brain has been observed at 4 T (20, 26) and was even more pronounced at 7 T (26). SNR measured in the center of the central transaxial slice of the head was higher by 30% at 4 T (20, 26) and by 75% at 7 T (26) as compared to the peripheral SNR. Attempts to compensate the RF transmit field inhomogeneity both numerically (29, 30) and experimentally (31) have been made using multi-port excitation of a single volume coil (29) or of a tranceive phased array (30, 31). These methods improve field homogeneity due to combining a homogeneous mode with higher order modes produced by multi-mode volume coils or tranceive phased arrays, but result in substantial phase distortion. Thus, there remains a need for improved radio frequency (RF) field homogeneity and signal-to-noise ratio (SNR) enhancement for body imaging, especially for imaging deeper regions such as central brain structures.